Ultrasound medical imaging is a commonly used procedure to produce images of internal body cavities such as blood vessels and surrounding tissue. In ultrasound imaging of a blood vessel, an Intravascular Ultrasound (“IVUS”) catheter is typically inserted into the blood vessel in a known manner. The IVUS catheter comprises an elongated member with an ultrasound transducer located at a distal end of the elongated member. The elongated member is inserted into the blood vessel, and the ultrasound transducer is positioned at a desired location in the blood vessel. The transducer emits ultrasound waves in the blood vessel or other such cavity when excited by a pulse. A portion of the emitted ultrasound waves is reflected back to the ultrasound transducer by tissue boundaries. The reflected ultrasound waves induce an echo signal at the ultrasound transducer. The echo signal is transmitted from the ultrasound transducer to an ultrasound console, which typically includes an ultrasound image processor, such as a computer, and a display. The display may comprise a monitor and/or a printer. The ultrasound console uses the received echo signal to image the cavity.
The echo signal is a serial amplitude modulated signal in which the amplitude of the signal varies with time. A typical echo signal has a time length of 8 μs, which corresponds to an image depth of approximately 6 millimeters from the ultrasound transducer. The echo signal carries both image brightness information and image depth information, where depth may be taken with respect to the ultrasound transducer. The image brightness information is provided by the amplitude of the echo signal. The image depth information is provided by the time position within the echo signal. An earlier time position in the echo signal corresponds to a lower image depth than a later time position in the echo signal.
In order to produce a radial cross-sectional image of a blood vessel and surrounding tissue, the ultrasound transducer is typically rotated along the axis of the elongated member. As the ultrasound transducer is rotated, the ultrasound transducer emits ultrasound waves in different radial directions. The resulting echo signals from the different radial directions are processed by the ultrasound console to produce a radial cross-sectional image of the blood vessel and the surrounding tissue. Alternatively, the ultrasonic transducer may be mounted in an assembly together with a reflective member (mirror), where the transducer emits ultrasonic energy in a substantially axial direction and the mirror is oriented to deflect the emitted ultrasonic energy in a radial direction.
Optical Coherence Tomography (“OCT”) is a type of optical coherence-domain reflectometry that uses low coherence interferometry to perform high resolution ranging and cross-sectional imaging. In OCT systems, a light beam from a low coherence light source is split into a reference light beam and a sample light beam. A diffraction grating may be used to provide a time delay. The sample light beam is directed onto a sample and the light scattered from the sample is combined with the reference light beam. The combination of the sample and reference light beams results in an interference pattern corresponding to the variation in the sample reflection with the depth of the sample, along the sample beam. The sample beam typically suffers a high loss of energy due to its interaction with the sample. The reference beam serves as a local oscillator to amplify the interference pattern to a detectable level and therefore must have a much higher energy level than the sample light beam. The interference pattern is detected by a photo detector, whose output is processed to generate a cross-sectional image of the sample. High resolution (less than 10 micrometer) imaging of the cross-sections of the sample by OCT is useful in biological and medical examinations and procedures, as well as in materials and manufacturing applications. An advantage of the above-described OCDR system is that the array of photo detectors is able to capture image brightness information at multiple image depths in one instance. This enables the OCT system to produce images at true video rates, e.g., 30 frames per second
OCT based systems may be implemented with fiber optics and an optical fiber carrying the sample light beam may be incorporated into a catheter or an endoscope for insertion into internal body cavities and organs, such as blood vessels, the gastrointestinal tract, the gynecological tract and the bladder, to generate images of internal cross-sections of the cavities or organs. The sample beam is typically emitted from the distal end of the instrument, where a prism or a mirror, for example, directs the sample light beam towards a wall of the cavity. The optical fiber and the prism or mirror may be rotated by a motor to facilitate examination of the circumference of the cavity.
An example of a fiber optic OCT system is shown in U.S. Pat. No. 5,943,133 (“the 133 patent”), where sample and reference light beams are carried in respective optical fibers to a diffraction grating, which introduces a time delay and also combines the sample and reference light beams. FIG. 1 is a schematic diagram of a system 10 disclosed in the '133 patent. The system includes a light source 12 optically coupled to a 50/50 beam splitter 14 through an optical fiber 16. The beam splitter 14 splits the incident light beam equally into a sample light beam and a reference light beam. The sample light beam is carried by an optical fiber 18 to a focusing lens 20, which focuses the sample light beam onto a sample 22. The optical fiber 18 may be contained within a catheter (not shown) for insertion into a body cavity, such as a blood vessel, for examination of the tissue of the wall of the cavity. Light received from the tissue is focused by the lens 20 and coupled back into the optical fiber. The received light travels back to the beam splitter 14, where it is split again. A portion of the received light is directed into another optical fiber 24, which conveys the light to a first collimator 26. The reference light beam travels through an optical fiber 28 to a second collimator 30. The first and second collimators 26, 30 direct the sample and reference light beams onto the same region of a diffraction grating 32. The diffracted, combined light beam is conjugated on the detector plane of a multi channel linear diode array detector 34 by a conjugating 36 lens. A neutral density filter (not shown) is provided to decrease the energy in the reference beam to prevent saturation of the detector.
The sample light beam suffers a significant loss of energy due to its interaction with the sample. The second pass through the 50/50 beam splitter further reduces the already attenuated light beam. In addition, the interaction of the light beams with the diffraction grating causes a further loss in both the sample light beam and the reference light beam of about 50% of the incident light in the first order. The diffraction grating also introduces noise. As a result, the system of the '133 patent has a low signal-to-noise ratio.
Another interferometric system using a diffraction grating is described in “Nonmechanical grating-generated scanning coherence microscopy”, Optics Letters, Vol. 23, No. 23, Dec. 1, 1998. FIG. 2 is a schematic diagram of the disclosed system 50. A light source 52 provides light to a 50/50 beam splitter 54 that splits the energy in the light beam equally into a sample light beam 55 and a reference light beam 56. The sample light beam 55 is directed to a focusing lens 58 that focuses the sample light beam onto a sample 60. The light received by the focusing lens 58 from the sample 60 is returned to the beam splitter 54. The reference light beam 56 is directed to a diffraction grating 62 in a littrow configuration. The diffracted reference light beam is also returned to the beam splitter 54. The sample and reference light beams are then combined in the beam splitter 54 and directed to a charge-coupled device (CCD) array 64 for detection and processing by a computer 66. The reference light beam needs to be suppressed here, as well.
Here, only the reference light beam is diffracted, making the system 50 more efficient than the system 10 of the '133 patent, shown in FIG. 1. However, the sample and reference arms in the system 50 of FIG. 2 cannot both be implemented with fiber optics. The diffraction grating introduces a time delay that is spatially spread across the width of the beam. The detector is a multi-element detector at least as wide as the light beam. Each element of the detector receives a portion of the beam corresponding to its position on the diffraction grating. If the reference light beam is conveyed by an optical fiber from the diffraction grating to the detector, the spatial order is lost. If the sample arm is implemented in fiber optics but the reference arm is not, the length of the reference arm would be inconveniently long.
It would be desirable to provide an imaging system that can process both ultrasound imaging data and optical interferometric imaging data with the same signal processor to display both ultrasound images and OCDR images, thereby reducing costs.